The present invention relates to non-invasive, real-time Doppler ultrasonic monitoring and feedback control of the extent and geometry of thermal damage in tissue treated by thermal therapy.
In recent years, thermal therapy using various modalities to thermally destroy benign and malignant lesions like tumors has gained widespread acceptance. The thermal modalities include lasers, electromagnetic wave, thermistors, and ultrasound. Thermal energy is delivered to the tissue of interest either externally or by interstitial means.
Hyperthermia is a popular thermal therapy for tumors and other tissue-related diseases. According to this method of treatment, a tumorous region inside a patient's body cavity is warmed to and kept at a temperature range of about 42.degree. to 50.degree. C. A number of methods of inducing hyperthermia have been tried, including electromagnetic wave (radiofrequency and microwave), whole body heating by external and extracorporeal means, and ultrasound. Hyperthermia has been applied both externally and by interstitial methods. However, with all these techniques, the problem exists of getting the energy to the target to cause the required cell death with predictability and precision, yet causing little or no damage to surrounding tissues. This problem is particularly pronounced with the heating of deep seated tumors, such as those in the liver.
Because they can be accurately focused and controlled, lasers in particular have become an accepted tool for thermal treatment (e.g., hyperthermia, tissue coagulation and ablation). Lasers are especially suited for thermal therapy, as they can deliver high energy directly into the tissue being treated so as to minimize the effects on surrounding normal tissue areas. The laser energy is typically applied to the tissue of interest through optical fibers. The optical fibers may be introduced percutaneously, via a blood vessel, through other body openings, or during surgical exposure of the tissue.
In addition to tissue heating and thermal injury, laser irradiation can lead to tissue vaporization, melting, ejection, and pyrolysis, which all result in removal of biological material. These effects have been collectively described as "tissue ablation," a process consisting of a cascade of events, each involving threshold dependent mechanisms. Tissue ablation can also result in charring and tearing of tissue.
In many thermal therapy procedures, the desire is to coagulate and kill the tissue rather than to ablate or vaporize it. In coagulation, the absorbed laser energy heats cells throughout the target volume to temperatures exceeding protein denaturation thresholds of approximately 65.degree. C. The resulting protein denaturation induces coagulative necrosis.
Interstitial laser photocoagulation (ILP) is a technique by which sufficient laser energy is deposited at low power levels so that thermal diffusion causes tissue coagulation while avoiding significant tissue vaporization near the fiber tips. Coagulative necrosis is believed to occur immediately during interstitial laser treatment.
Accurately predicting tissue thermal damage according to theories and computer modelling is very difficult, if not impossible, and is also unreliable. This is due to the heterogeneity of tissues with respect to: (1) physical properties (e.g., optical and thermal properties); (2) tissue geometry; and (3) blood perfusion in tissues. The transport of thermal energy in tissues is a complex process involving conduction, convection, radiation, metabolism, evaporation and phase change. Additionally, the broad parameter space of treatment operation (energy delivered, operation duration and delivery geometry, etc.) dramatically influence the overall tissue response. All these factors result in nonuniformly distributed and highly dynamic temperature fields in tissues during thermal treatment, yielding highly unpredictable tissue thermal damage.
Thus, one of the major obstacles to effective and wide-spread applications of thermal therapies has been the lack of a reliable noninvasive and real-time detection method to determine the extent and geometry of tissue damage. Such techniques can not only improve the optimization of the treatment parameters, but can also provide means for protecting critical organs surrounding treated tissues, resulting in optimal treatment.
Although laser-induced damage extends deeply into the tissue, the visible effect to the naked eye is superficial. Therefore, the use of medical imaging apparatus has proved useful to evaluate tissue response during thermal therapy. Indeed, imaging laser-induced damage is currently a major challenge in the field of laser medicine. The imaging modalities presently available for the acquisition of clinical images include two-dimensional X-ray imaging, computed tomographic imaging (CT), magnetic resonance imaging (MRI), ultrasonography, two-dimensional radioisotope imaging, single photon emission computed tomography, positron emission tomography, thermography, and transillumination. Several of these methods have been proposed for measuring temperature and detecting tissue response during laser therapy, including implanted thermistors or thermocouples, CT imaging, MRI, and ultrasonic imaging.
Methods such as thermocouple or light-detector insertion can provide information about the light distribution or heat development at different points in the tissue. These parameters may be used as feedback tools for laser adjustment during therapy to achieve optimal localized tumor destruction. However, the effectiveness of such probes is limited because the probes have to be placed invasively into the tissue, and their position is rather critical.
The use of magnetic resonance imaging (MRI) has also been investigated as a means of visualizing tissue response during laser treatment. However, currently available MRI devices require a relatively long time to collect data for an image, and thus may not be used practically for real-time evaluation. MRI devices are also large and expensive.
Ultrasound imaging is another technique that has gained favor of late for use in conjunction with thermal therapy. In medical ultrasound imaging, pulses of longitudinal sound waves at frequencies from 1-20 MHz are emitted by one or more piezoelectric transducers into the body volume being imaged. Inside the body, ultrasound is attenuated through scattering (including reflection, refraction and diffraction) and absorption. The intensities and arrival times of ultrasound echoes, that is, of waves reflected back to the transducer(s) by internal acoustic boundaries, are measured and converted into images of the reflecting boundaries. For sound waves, a boundary is a spatial discontinuity in the acoustic impedance, defined in any medium as the product of the speed of sound and density. The speed of sound and acoustic impedance are temperature dependent.
Several investigators have reported expanding hyperechoic regions in ultrasonic images made during laser irradiation of tissues. For example, hyperechoic regions have been reported in pig liver during and after laser irradiation. See Malone, et al., "Sonographic Changes During Hepatic Interstitial Laser Photocoagulation: An Investigation of Three Optical Fiber Tips," Investigative Radiology, vol. 27, pp. 804-13 (Oct. 1992); Dachman et al., "US-Guided Percutaneous Laser Ablation of Laser Tissue in a Chronic Pig Model," Radiology, vol. 176, pp. 129-33 (1990). In some cases, it was reported that a growing hyperechoic region was found during irradiation which sometimes became hypoechoic after one hour. See Godlewski, et al., "Ultrasonic and Histopathological Correlations of Deep Focal Hepatic Lesions Induced by Stereotaxic Nd-YAG Laser Applications," Ultrasound in Med. & Biol., vol. 14, pp. 287-91 (1988); Godlewski, et al., "Deep Localized Neodymium (Nd)-YAG Laser Photocoagulation in Liver Using a New Water Cooled and Echoguided Handpiece," Lasers in Surgery and Medicine, vol. 8, pp. 501-09 (1988). Hypoechoic regions have also been found in rat liver after laser treatment of tumors. Van Hillegersberg, et al., "Water-Jet-Cooled Nd:YAG Laser Coagulation of Experimental Liver Metastases: Correlation Between Ultrasonography and Histology," Lasers in Surgery and Medicine, vol. 13, pp. 332-43 (1993). Similarly, hypoechoic regions have been found in canine myocardium irradiated with a low power laser. Watanabe, et at., "Thermally Controlled Laser Irradiation of the Myocardium with Intraoperative Ultrasound Monitoring," PACE, vol. 13, pp. 653-62 (May 1990).
Because of the large increase in echogenicity which subsequently decreases when the laser is turned off, it has been postulated that vapor or microbubble formation is responsible for the increase. Residual echogenicity is probably related to changes in the structure and to reorganization of tissue during coagulation and ablation.
Air pockets created during interstitial laser photocoagulation, either as small bubbles or a cavity at the fiber tip, will generate strong ultrasound echoes. It is thus possible to obtain ultrasound contrast images of thermal lesions induced by thermal therapy. Further, the use of imaging systems for potential control of interstitial laser photocoagulation was recently reviewed, with mixed results. See Wyman et al., "Medical Imaging Systems for Feedback Control for Interstitial Laser Photocoagulation," Proceedings of the IEEE, vol. 80, p. 890-902 (June 1992). Both MRI and ultrasound were able to detect changes in tissue properties during and after laser irradiation. The results with ultrasound, however, were inconsistent. It was postulated that actual tissue vaporization (ablation) may be required to visually observe substantial and consistent increases in echogenicity. In addition, it appears that respiratory motion poses a problem for ultrasound image-based control of interstitial laser therapy.
In addition to enabling visual evaluation of the extent of thermal tissue damage, ultrasound has also been used to guide placement of a laser. See Roth, R. and Aretz, H., "Transurethral Ultrasound-Guided Laser-Induced Prostatectomy (Tulip Procedure): A Canine Prostate Feasibility Study," The Journal of Urology, vol. 146, pp. 1128-35 (Oct. 1991). Therein, a combined ultrasound/laser device is described for transurethral laser coagulation of the prostate. Ultrasonic images were used for guidance and also for evaluating the results of laser treatment in dogs. At higher laser power levels (i.e., greater than 50 watts), hyperechoic regions correlating to the area of coagulation necrosis were seen.
Laser ablation of atherosclerotic plaques is another area where real-time monitoring is needed. See Borst, et al., "Laser Ablation and the Need for Intra-Arterial Imaging," International Journal of Cardiac Imaging, vol. 4, pp. 127-33 (1989). Several investigators have used ultrasound guidance for laser ablation of plaques with mixed results. For example, in Aretz, et al., "Intraluminal Guidance of Transverse Laser Coronary Atherectomy," International Journal of Cardiac Imaging, vol. 4, pp. 153-57 (1989), the concept of a combined laser and atherectomy and ultrasonic imaging catheter is described. This device was subsequently evaluated for imaging (but not lasing) in dogs. Aretz, et al., "Ultrasound Guidance of Laser Atherectomy," International Journal of Cardiac Imaging, vol. 6, pp. 231-37 (1991). Also, in Duda et al., "Ultrasound-Monitored Laser Angioplasty: Preliminary Clinical Results," Cardiovasc. Intervent. Radiol., vol. 16, pp. 89-92 (1993), a clinical use of a combined laser angioplasty and ultrasonic imaging catheter to treat femoral and iliac stenoses was reported. The ultrasound image was found not to be useful in directing laser ablation, but it was found useful to measure the residual stenosis after treatment. Cavitation bubbles were seen in the blood during irradiation, but no changes in the echogenicity of the vessel wall were noticed after laser treatment. A major problem reported was that the ultrasound beam scanned radially while the laser beam was axial to the catheter, thereby preventing real-time monitoring.
An ultrasonic displacement measuring system based on Doppler ultrasound principles has been described for detecting myocardial thickening. See Hartley, et al., "An Ultrasonic Method for Measuring Tissue Displacement: Technical Details and Validation for Measuring Myocardial Thickening," IEEE Transactions on Biomedical Engineering, vol. 38, pp. 735-747 (Aug. 1991), the disclosure of which is herein incorporated by reference.
The technique is based on principles of both pulse-echo and pulsed Doppler ultrasound. By detecting and following the phase of the echoes returned from the transducer, the motion of the reflector with respect to the transducer can be quantified. The instantaneous phase of the returning echo is proportional to the distance of the reflector from the transducer. See Hartley et al. (1991) at p. 736. Using a tone burst mode of operation, the transmitted signal (S.sub.t) has the following form: ##EQU1## where .omega. is the angular frequency of the transmitted wave, T is the pulse repetition period which must be an integer number of cycles of .omega., n is a positive integer, and t.sub.x is the duration of the transmitted burst. This tone burst is propagated from the transducer toward a target where part of it is reflected back toward the transducer which then acts as a receiver.
The received echo signal (S.sub.r) has the form ##EQU2## where a is the amplitude of the received echo signal and t.sub.d is the time delay between the beginning of the transmitted burst and the beginning of the received echo signal. In turn, EQU t.sub.d =2d/c (3)
where d is the distance from the transducer to the reflecting interface and c is the speed of sound in the target. Rearranging and substituting results in: ##EQU3## Since .omega.=2.pi.c/.lambda., eq. (4) can be put in the form ##EQU4## where .lambda. is the wavelength of the ultrasonic wave in the conducting medium (such as tissue), and where EQU .phi.=4.pi.d/.lambda.. (6)
In eq. (6), .phi. represents the phase (in radians) of the echo signal with respect to the transmitter signal and is directly proportional to the distance d from the transducer to the reflecting target.
A block diagram of a prior art ultrasonic displacement measuring instrument is shown in FIG. 1. In the instrument described in FIG. 1, phase is sensed by a quadrature-phase detector 32 consisting of two analog multipliers. The reference inputs to the multipliers are cos (.omega.t) [27] and sin (.omega.t) [29] and are derived from a master oscillator 26 which runs continuously at angular frequency .omega.. Multiplying each of these signals by the echo signal results in EQU cos (.omega.t).times.a cos (.omega.t-.phi.)=a{cos (2.omega.t-.phi.)+cos (.phi.)}/2 (7)
and EQU sin (.omega.t).times.a cos (.omega.t-.phi.)=a{sin (2.omega.t-.phi.)+sin (.phi.)}/2 (8)
for t as in eq. (2).
If the above signals are low-pass filtered to remove the high frequency terms at 2.omega.t, are multiplied by 2 to eliminate the 1/2, and are sampled only during the received interval [i.e., t as in eq. (2)], the sampled, phase detected signals then become EQU x=a cos .phi. (9)
and EQU y=a sin .phi., (10)
where x and y can be considered as components of a polar coordinate phase vector of length a and angle .phi.. Components x and y are called "quadrature range-phase signals."
In general, the target will be moving with respect to the transducer, which will generate a Doppler shift and cause a rotation of the phase vector. To see how the phase and Doppler shift frequency are related, consider a target moving at velocity v at an angle .theta. with respect to the sound beam axis. Its distance from the transducer d is given by EQU d=.intg..sub.o.sup.t v cos .theta.dt=vt cos .theta.+d.sub.0( 11)
where d.sub.0 is the initial position of the target. If eq. (11) is substituted into eq. (6), phase also becomes a function of time given by EQU .phi.=(4.pi.vt/.lambda.) cos .theta.+constant. (12)
Differentiating the phase yields an angular frequency (.omega..sub.d): EQU .omega..sub.d =d.phi./dt=(4.pi.vt/.lambda.) cos .theta.. (13)
Substituting .omega..sub.d =2.pi.f.sub.d and .lambda.=c/f, eq. (13) becomes EQU f.sub.d =(2fV/c) cos .theta., (14)
which is the well-known Doppler equation with f.sub.d being the Doppler shift frequency, in Hz, of the reflected wave. The Doppler shift can thus be obtained from the phase by differentiation. For a moving target, the phase vector rotates with an angular frequency given by .omega..sub.d in a direction corresponding to the direction of the target motion: clockwise for motion away from the transducer (receding phase) or counterclockwise for motion toward the transducer (advancing phase).
Timing for the instrument shown in FIG. 1 (operated at 10 MHz) is controlled by 10 MHz crystal oscillator 26. The 10 MHz frequency is divided by 2560 by frequency divider 14 to produce a pulse repetition frequency (PRF) of 3.90625 kHz. A 0.4 .mu.s pulse from pulse generator 12 is used to gate 4 cycle bursts of the 10 MHz signal [see eq. (1)] to transmitter amplifier 28, which drives ultrasonic transducer 10. Transducer 10 converts the electrical signals to acoustic tone bursts, which are propagated into the tissue where they are reflected by structures along the sound beam. The echoes returning to transducer 10 are converted back into electrical signals [see eq. (2)], which are amplified by RF amplifier 30 to produce signal 31 (see eq. (4)) and compared by quadrature-phase detector 32 in phase to quadrature signals {cos (.omega.t) [27] and sin (.omega.t) [29]} from 10 MHz oscillator 26. The two phase detector outputs 33 and 35 (which are quadrature signals that correspond to eq. (7) and (8), respectively) are then sampled by dual sample and hold circuit 34 with a 0.2 .mu.s range-gate pulse 24 delayed by 2-50 .mu.s from the transmit pulse by variable delay circuit 18. The range of circuit 24 is selected by potentiometer 16. After sampling, the two signals are high-pass filtered at 1 Hz to remove the dc components from the stationary structures and low-pass filtered at 1 kHz by dual filters 36 to remove residual signals. Except for the lower bandwidth, the sampled, filtered signals 37 and 39 are the in-phase (I) and quadrature-phase (Q) Doppler signals (which correspond to eq. (9) and (10), respectively). Signals 37 and 39 resemble quadrature audio signals from a pulsed Doppler instrument for measuring blood flow, and may be received at output 52.
The vector representation of the quadrature signals may be shown in X-Y display 54 shown in FIG. 1. The radius a represents the amplitude of the echo from the target, and the phase .phi. of the echo represents the position of the target. The change in position (or displacement) can be measured by noting the direction (clockwise or counterclockwise) and counting the revolutions of the vector. Each revolution corresponds to reflector motion of 0.075 mm at 10 MHz. To improve the resolution of the instrument to 0.019 mm, revolutions are counted in 90.degree. increments corresponding to axis crossings in the X-Y display 54. Logic circuits in up-down counter controller 40 detect positive and negative zero crossings of each quadrature signal, assign a direction based on the polarity of the zero crossing and the polarity of the other signal at the time, and increment or decrement 8-bit up-down counter 40. Full-scale range for the 8-bit counter shown is about 4.8 mm. Digital-to-analog converter (DAC) 44 receives 8-bit output 42 from counter 40 and produces a voltage output 46 that represents the change in position of echoes within the sample volume with a calibration of 2 V/mm. Since no filter is used on displacement output 46, it is updated after each sample (approximately 4 kHz) whenever the reflector has moved 0.019 mm.
Signals available from the instrument shown in FIG. 1 include: displacement 46 at 2 V/mm, analog range 20 at 0.1 V/cm, quadrature signals 52, and quadrature audio 52a from audio amplifier 50 and speaker 48. In addition, monitor outputs 21, 25, and 33 for oscilloscope 22 are provided from transmit pulse generator 12 (for triggering), range-gate pulse generator 24, and phase detector 32, respectively. Controls are a 2-40 mm range-gate potentiometer 16 and a 6, 7, or 8 bit limit switch 41 to up/down counter 40. Inputs are from a 10 MHz piezoelectric transducer 10 attached or planted within the tissue of interest, and a reset command 55 from a triggered event.
No provision is made in the prior art circuit of FIG. 1, however, for real-time evaluation of the signals received by transducer 10 in relation to tissue thermal response, nor of controlling the extent or geometry of tissue thermal damage resulting from thermal treatment of a lesion in living tissue. Thus, a need exists for an apparatus and method for noninvasive, real-time monitoring and control of the extent and geometry of tissue damage in thermal therapies.